Optical coherence tomography (OCT) is a noninvasive imaging technique that provides microscopic tomographic sectioning of biological samples. By measuring singly backscattered light as a function of depth, OCT fills a valuable niche in imaging of tissue ultrastructure, providing subsurface imaging with high spatial resolution (˜2.0-10.0 μm) in three dimensions and high sensitivity (>110 dB) in vivo with no contact needed between the probe and the tissue.
In biological and biomedical imaging applications, OCT allows for micrometer-scale imaging non-invasively in transparent, translucent, and/or highly-scattering biological tissues. The longitudinal ranging capability of OCT is generally based on low-coherence interferometry, in which light from a broadband source is split between illuminating the sample of interest and a reference path. The interference pattern of light reflected or backscattered from the sample and light from the reference delay contains information about the location and scattering amplitude of the scatterers in the sample. In time-domain OCT (TDOCT), this information is typically extracted by scanning the reference path delay and detecting the resulting interferogram pattern as a function of that delay. The envelope of the interferogram pattern thus detected represents a map of the reflectivity of the sample versus depth, generally called an A-scan, with depth resolution given by the coherence length of the source. In OCT systems, multiple A-scans are typically acquired while the sample beam is scanned laterally across the tissue surface, building up a two-dimensional map of reflectivity versus depth and lateral extent typically called a B-scan. The lateral resolution of the B-scan is approximated by the confocal resolving power of the sample arm optical system, which is usually given by the size of the focused optical spot in the tissue.
The time-domain approach used in conventional OCT, including commercial instruments, such as Carl Zeiss Meditec's Stratus® and Visante® products, has been successful in supporting biological and medical applications, and numerous in vivo human clinical trials of OCT reported to date have utilized this approach.
An alternate approach to data collection in OCT has been shown to have significant advantages in increased signal-to-noise ratio (SNR). This approach involves acquiring the interferometric signal generated by mixing sample light with reference light at a fixed group delay as a function of optical wavenumber. Two distinct methods have been developed which use this Fourier domain OCT (FD-OCT) approach. The first, generally termed Spectral-domain or spectrometer-based OCT (SDOCT), uses a broadband light source and achieves spectral discrimination with a dispersive spectrometer in the detector arm. The second, generally termed swept-source OCT (SSOCT) or optical frequency-domain imaging (OFDI), time-encodes wavenumber by rapidly tuning a narrowband source through a broad optical bandwidth. Both of these techniques may allow for a dramatic improvement in SNR of up to 15.0-20.0 dB over time-domain OCT, because they typically capture the A-scan data in parallel. This is in contrast to previous-generation time-domain OCT, where destructive interference is typically used to isolate the interferometric signal from only one depth at a time as the reference delay is scanned.
FDOCT systems are discussed below with respect to FIGS. 1 through 3. Referring first to FIG. 1, a block diagram illustrating a Fourier domain OCT system in accordance with some embodiments of the present inventive concept will be discussed. As illustrated in FIG. 1, the system includes a broadband source 100, a reference arm 110 and a sample arm 140 coupled to each other by a beamsplitter 120. The beamsplitter 120 may be, for example, a fiber optic coupler or a bulk or micro-optic coupler without departing from the scope of the present inventive concept. The beamsplitter 120 may provide from about a 50/50 to about a 90/10 split ratio. As further illustrated in FIG. 1, the beamsplitter 120 is also coupled to a wavelength or frequency sampled detection module 130 over a detection path 106 that may be provided by an optical fiber.
As further illustrated in FIG. 1, the source 100 is coupled to the beamsplitter 120 by a source path 105. The source 100 may be, for example, a SLED or tunable source. The reference arm 110 is coupled to the beamsplitter over a reference arm path 107. Similarly, the sample arm 140 is coupled to the beamsplitter 120 over the sample arm path 108. The source path 105, the reference arm path 107 and the sample arm path 108 may all be provided by optical fiber.
As further illustrated in FIG. 1, the sample arm 140 may include scanning delivery optics and focal optics 160. Also illustrated in FIG. 1 is the reference plane 150 and a representation of an OCT imaging window 170.
Referring now to FIG. 2, a block diagram of an FDOCT retinal imaging system will be discussed. As illustrated in FIG. 2, in an FDOCT retinal imaging system, the reference arm 110 may further include a collimator assembly 280, a variable attenuator 281 that can be neutral density or variable aperture, a mirror assembly 282, a reference arm variable path length adjustment 283 and a path length matching position 250, i.e. optical path length reference to sample. As further illustrated, the sample arm 240 may include a dual-axis scanner assembly 290 and a variable focus objective lens 291.
The sample in FIG. 2 is an eye including a cornea 295, iris/pupil 294, ocular lens 293 and retina 296. A representation of an OCT imaging window 270 is illustrated near the retina 296. The retinal imaging system relies in the optics of the subject eye, notably cornea 295 and ocular lens 293, to image the posterior structures of the eye.
Referring now to FIG. 3A, a block diagram illustrating a FDOCT cornea imaging system will be discussed. As illustrated therein, the system of FIG. 3A is very similar to the system of FIG. 2. However, the objective lens variable focus need not be included, and is not included in FIG. 3A. The anterior imaging system of FIG. 3A images the anterior structures directly, without reliance on the optics of the subject to focus on the anterior structures.
As illustrated by FIGS. 3A through 3C, the OCT imaging window 370 can be moved to image various portions of the sample.
In both spectrometer-based and swept-source implementations of FDOCT, light returning from all depths is generally collected simultaneously, and is manifested as modulations in the detected spectrum. Transformation of the detected spectrum from wavelength to wavenumber (or frequency), correction for dispersion mismatches between the sample and reference arms, and Fast Fourier transformation typically provides the spatial domain signal or “A-scan” representing depth-resolved reflectivity of the sample. The uncorrected A-scan may also include a strong DC component at zero pathlength offset, so-called “autocorrelation” artifacts resulting from mutual interference between internal sample reflections, as well as both positive and negative frequency components of the depth-dependent cosine frequency interference terms. Because of this, FDOCT systems typically exhibit a “complex conjugate artifact” due to the fact that the Fourier transform of a real signal, the detected spectral interferogram, is typically Hermitian symmetric, i.e., positive and negative spatial frequencies are not independent. As a consequence, sample reflections at a positive displacement, relative to the reference delay, typically cannot be distinguished from reflections at the same negative displacement, and appear as upside-down, overlapping images on top of genuine sample structure, which generally cannot be removed by image processing.
The maximum single-sided imaging depth available in SDOCT is governed by the spectral sampling interval. The maximum single-sided imaging depth is inversely proportional to the spectral sampling interval. With a fixed number of sampled spectral elements, there is an inverse relationship between the maximum imaging depth and the minimum axial resolution of the imaging system. In commercial FDOCT systems at 830 nm and 1300 nm reported to date, the single-sided imaging depth has been limited to approximately 4 mm. Time domain imaging has been used for greater imaging depths.
The finite spectral resolution of any real FDOCT system, whether governed by the linewidth of a swept laser source in SSOCT, or the geometric optical performance of the spectrometer convolved with the finite pixel size of the detector array in SDOCT, gives rise to a sensitivity “falloff” with imaging depth into the sample. It is common to have greater than 6 dB degradation in signal-to-noise from the position of zero reference delay to the position of maximum single-sided depth. This sensitivity “falloff” limits the portion of the single-sided depth useful for imaging.
To reduce the impact of these limitations in FDOCT imaging, imaging is commonly performed with the entire sample either above or below the reference position, limiting the available imaging depth to between 2 mm and 4 mm, and placing the sample region of interest close to the zero reference delay position.
Each of these constraints poses limitations on the application of FDOCT to clinical ophthalmology. Imaging systems have generally been dedicated to imaging of specific anatomy, such as retina or cornea, where the mirror image artifacts do not fold over onto images of the region of interest. Utility to image deeper anatomic structures, such as the choroid, has been limited by sensitivity “falloff”.
Addressing these limitations opens significant new application areas for FDOCT, particularly in ophthalmology. Full range volumetric anterior segment imaging (cornea to lens) for improved diagnosis of narrow angle glaucoma is enabled at speeds 20 times greater and resolutions four times finer than time domain implementations. Real-time image guided surgery, for anterior chamber, cataract, or retina, is enabled by allowing placement of a deep imaging window at any position within the sample, without concern for confounding mirror image artifacts or signal “falloff.” Images of the entire eye may be acquired, enabling for the first time modeling in three dimensions the entire optical structure of the eye and enabling whole-eye biometry.